Implantable stimulation devices are devices that generate and deliver electrical stimuli to body nerves and tissues for the therapy of various biological disorders, such as pacemakers to treat cardiac arrhythmia, defibrillators to treat cardiac fibrillation, cochlear stimulators to treat deafness, retinal stimulators to treat blindness, muscle stimulators to produce coordinated limb movement, spinal cord stimulators to treat chronic pain, cortical and deep brain stimulators to treat motor and psychological disorders, and other neural stimulators to treat urinary incontinence, sleep apnea, shoulder sublaxation, etc. The present invention may find applicability in all such applications, although the description that follows will generally focus on the use of the invention within a Spinal Cord Stimulation (SCS) system, such as that disclosed in U.S. Pat. No. 6,516,227, which is incorporated herein by reference in its entirety.
Spinal cord stimulation is a well-accepted clinical method for reducing pain in certain populations of patients. As shown in FIGS. 1A and 1B, a SCS system typically includes an Implantable Pulse Generator (IPG) 100, which includes a biocompatible case 30 formed of titanium for example. The case 30 typically holds the circuitry and power source or battery necessary for the IPG to function, although IPGs can also be powered via external RF energy and without a battery. The IPG 100 is coupled to electrodes 106 via one or more electrode leads (two such leads 102 and 104 are shown), such that the electrodes 106 form an electrode array 110. The electrodes 106 are carried on a flexible body 108, which also houses the individual signal wires 112 and 114 coupled to each electrode. In the illustrated embodiment, there are eight electrodes on lead 102, labeled E1-E8, and eight electrodes on lead 104, labeled E9-E16, although the number of leads and electrodes is application specific and therefore can vary.
As shown in FIG. 2, the IPG 100 typically includes an electronic substrate assembly 14 including a printed circuit board (PCB) 16, along with various electronic components 20, such as microprocessors, integrated circuits, and capacitors mounted to the PCB 16. Two coils are generally present in the IPG 100: a telemetry coil 13 used to transmit/receive data to/from an external controller 12 as explained further below; and a charging coil 18 for charging or recharging the IPG's power source or battery 26 using an external charger (not shown). The telemetry coil 13 can be mounted within the header connector 36 as shown.
As just noted, an external controller 12, such as a hand-held programmer or a clinician's programmer, is used to send data to and receive data from the IPG 100. For example, the external controller 12 can send programming data to the IPG 100 to dictate the therapy the IPG 100 will provide to the patient. Also, the external controller 12 can act as a receiver of data from the IPG 100, such as various data reporting on the IPG's status. The external controller 12, like the IPG 100, also contains a PCB 70 on which electronic components 72 are placed to control operation of the external controller 12. A user interface 74 similar to that used for a computer, cell phone, or other hand held electronic device, and including touchable buttons and a display for example, allows a patient or clinician to operate the external controller 12.
Wireless data transfer between the IPG 100 and the external controller 12 takes place via inductive coupling. To implement such functionality, both the IPG 100 and the external controller 12 have coils 13 and 17 respectively. Either coil can act as the transmitter or the receiver, thus allowing for two-way communication between the two devices. When data is to be sent from the external controller 12 to the IPG 100 for example, coil 17 is energized with alternating current (AC), which generates a magnetic field 29, which in turn induces a voltage in the IPG's telemetry coil 13. The power used to energize the coil 17 can come from a battery 76, which like the IPG's battery 26 is preferably rechargeable, but power may also come from plugging the external controller 12 into a wall outlet plug (not shown), etc. The induced voltage in coil 13 can then be transformed at the IPG 100 back into the telemetered data signals. To improve the magnetic flux density, and hence the efficiency of the data transfer, the IPG's telemetry coil 13 may be wrapped around a ferrite core 13′.
As is well known, inductive transmission of data from coil 17 to coil 13 can occur transcutaneously, i.e., through the patient's tissue 25, making it particular useful in a medical implantable device system. During the transmission of data, the coils 13 and 17 lie in planes that are preferably parallel. Such an orientation between the coils 13 and 17 will generally improve the coupling between them, but deviation from ideal orientations can still result in suitably reliable data transfer.
To communicate a serial stream of digital data bits via inductive coupling, some form of modulation is generally employed. In a preferred embodiment, Frequency Shift Keying (FSK) can be employed, in which the logic state of a bit (either a logic ‘0’ or a logic ‘1’) corresponds to the frequency of the induced magnetic field 29 at a given point in time. Typically, this field has a center frequency (e.g., fc=125 kHz), and logic ‘0’ and ‘1’ signals comprise offsets from that center frequency (e.g., f0=121 kHz and f1=129 kHz respectively). Once the data is modulated in this manner at the transmitting device (e.g., the external controller 12), it is then demodulated at the receiving device (e.g., the IPG 100) to recover the original data. While FSK modulation may be preferred for a given application, one skilled in the art will recognize that other forms of data modulation (e.g., amplitude modulation, On-Off-Keying (OOK), etc.) can be used as well. These modulation schemes as used in a medical implantable device system are disclosed in U.S. Pat. No. 7,177,698, which is incorporated herein by reference in its entirety, and because they are well known, they are not further discussed.
Typical transceiver circuits 150 and 151 for effecting the transmission and reception of data in the manners just described are shown in FIGS. 3A and 3B. In the example shown, it is assumed that the transceiver circuits 150 and 151 are within the IPG 100, although it should be remembered that such circuitry may also be present in the external controller 12. Each circuit comprises a transmitter (TX), an L-C resonant circuit (or as it is sometimes known in the art, a “tank circuit”), and a receiver (RX). In both cases, the inductor (L) in the tank circuit comprises the IPG's data communication coil 13 discussed previously. In circuit 150 (FIG. 3A), the inductor L and capacitor C are connected in series; in circuit 151 (FIG. 3B), the inductor L and capacitor C are in parallel.
In either case, transmission and reception is effected in essentially the same way. As shown in the example of FIG. 3A, the transmitter TX modulates a digital input, D_tx, to produce two complementary drive signals (drive and drive′), which are applied across the resonant circuit. The digital input is referenced to the basic digital power supply operating in the IPG, Vdd, which may be about 2.8V. The magnitude of the drive signals, by contrast, are referenced to Vbat, i.e., the voltage of the battery 26 in the IPG 100, which may be about 3.0 to 4.2V. (Usually the power supply voltage, Vdd, is derived by a regulator from the battery voltage, Vbat, but this is not strictly necessary, and these voltages can be one and the same, and either can power either the transmitter or the receiver. For the purpose of this disclosure, either Vdd or Vbat may be considered as a power supply voltage). Because the drive signals are complimentary, +Vbat and ground are alternatively applied across the resonant circuit, causing the desired resonance to produce the magnetic field 29. When receiving, the receiver RX receives differential inputs caused by the resonance of the resonance circuit, which is then demodulated to form the digital output D_rx, which is again referenced to Vdd (or Vbat, again, either of these voltages can be considered the power supply voltage that powers the receiver). The circuitry for transmitters TX and receivers RX are well known, and hence are not further discussed.
Each of these series and parallel tank circuits has advantages and disadvantages. For example, the series-connected L-C tank 150 is capable of forming large voltages across the inductor, L during transmission. In other words, the voltage produced at the node between the inductor and the capacitor, VA, is amplified by the Q (quality factor) of the tank which can equal about +/−50V or so. This improves the magnitude of the magnetic field 29 which is produced, and thus ultimately improves the transmitter performance. As a result, a low voltage drive transmitter 160 can be used that drives the resonant circuit with smaller voltage signals compatible with standard CMOS integrated circuit technology. By contrast, the receiver RX in the series configuration is generally desired to have a relatively low input impedance 164 (e.g., <10 ohms) to enhance detection of the voltage induced in the resonant circuit by the received magnetic field 29. Unfortunately, the simultaneous desires for a high transmit field and low receiver input impedance increases the power consumption in the receiver RX. Increased receiver power consumption in the IPG 100 is especially problematic due when one considers that IPG batteries 26 (FIG. 2) are relatively small and therefore of limited capacity.
By contrast, the transmit field in the parallel-connected transceiver circuit 151 is not as high, because the voltages across the inductor are limited to the magnitude of the drive signals. As a result, a high drive transmitter 162 is required, which requires drive signals of greater magnitude (+/−50V or so), and which is not compatible with standard CMOS integrated circuit technology. However, the benefit to the parallel configuration occurs on the receiver side. Specifically, the receiver can have a relatively high input impedance 166 (e.g., >10 k ohms) compared to the low impedance receiver 164 used in the series configuration, resulting in lower power consumption and increased detection sensitivity in the receiver.